The field of the invention is nuclear magnetic resonance imaging (xe2x80x9cMRIxe2x80x9d) methods and systems. More particularly, the invention relates to the production of magnetic field gradients in MRI systems and their use in imaging pulse sequences.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or xe2x80x9ctippedxe2x80x9d, into the x-y plane to produce a net transverse magnetic moment Mt. An NMR signal is emitted by the excited spins after the excitation signal B1 is terminated and, this signal may be received and processed to form an image.
When utilizing NMR signals to produce images a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The region of interest may be a small portion of a patient""s anatomy, such as the head or heart, or a much larger portion, such as the entire thorax or spine. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well-known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (GXx, Gy, and Gz,) which have the same direction as the polarizing field B0, but which have a gradient along the respective X, Y and Z axes. The magnetic field gradients are produced by a trio of coil assemblies placed around the object being imaged. By controlling the strength of these gradients during each NMR measurement cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified.
In order to accommodate the imaging of large portions of a patient, each gradient field coil must produce a magnet field that varies linearly along one axis of a very large volume. On the other hand, to image a small portion of a patient, each gradient field coil may be smaller in physical size and have entirely different electrical characteristics than the larger whole-body gradient coils.
There are many conflicting design considerations when providing an optimal gradient subsystem for an MRI system. Factors such as peak gradient amplitude, peak gradient slew rate, gradient spatial linearity over the imaging volume, heat generation and patient safety with respect to peripheral nerve stimulation (i.e. dB/dt limits) all must be considered. Compromises must be made. For example, a very linear field gradient coil tends to be high in inductance and requires a high voltage power supply to provide a high slew rate. A coil which produces a linear gradient field over a large field of view for spinal imaging can produce excessive dB/dt when driven at a high slew rate. The efficiency (and hence heat generation) is better when a smaller gradient coil having a reduced diameter or shorter length is employed. Smaller diameter gradient coils are used, for example, when imaging the head with fast EPI pulse sequences.
A number of solutions have been proposed to address this gradient subsystem design dilemma. As described in U.S. Pat. No. 5,311,135, for example, the gradient coil windings may be tapped and the gradient amplifiers can be switched to different taps on the coils depending on the particular scan being conducted. Such switching changes the size and location of the optimal gradient fields as well as the electrical characteristics of the coils. In another solution described in U.S. Pat. No. 5,736,858, two separate sets of gradient coils are provided and the three gradient amplifiers may be switched to either or both sets depending on the particular scan being performed. One set of gradient coils is a relatively large, whole-body coil and a supplementary gradient coil set is relatively small. The gradient amplifiers have fast semiconductor switches capable of rapidly switching between three gradient coil configurations within a pulse sequence.
In yet another solution disclosed in co-pending U.S. patent application Ser. No. 09/382,905 filed on Aug. 25, 1999 and entitled xe2x80x9cModular Gradient System For MRI Systemxe2x80x9d separate sets of gradient amplifiers are used to drive the whole-body gradient coils and a supplementary gradient coils set. Both gradient coil sets can thus be employed separately or at the same time during an imaging pulse sequence.
For a number of imaging applications the supplementary gradient coil set is used because they are specifically designed for the application. Such supplementary gradient coil sets typically have a limited field of view over which their gradient fields are uniform. This can be a disadvantage for applications which require spatial localization near the edge or outside of the imaging volume. For example, spatial saturation as described in U.S. Pat. No. 4,715,383 is often applied near the edge of the field of view. If the Supplementary coil is used for both spatial saturation and imaging gradients, the spatial saturation pulses will only be effective over a small, limited volume.
In another clinical application, Fast Spin Echo (FSE) sequences are used to image spines with sagittal planes. Phase encoding is typically S/I to prevent cardiac, flow and respiratory motion artifacts from overlaying the spine image. Because the patient usually extends outside the field of view, phase encode oversampling, which doubles the field of view (xe2x80x9cNo Phase Wrapxe2x80x9d), is often used to avoid aliasing or wrap-around artifacts in the phase encoding direction. However, there is an artifact which is aliased into the field of view even when phase encode oversampling is used. This artifact is caused by a combination of B0 field inhomogeneity and gradient nonlinearity in the region beyond the normal imaging volume. Typically, the B1 field is strong enough in this area to flip spins, but insufficient gradient amplitude is available to properly frequency and phase encode the spins. The artifact is a bright spot which may be spread into a ghosted pattern because of system instabilities, eddy currents, and echo amplitude modulations which are typical in a FSE pulse sequence. Specifically this artifact will be highly visible when the supplementary coil set is used. Spatial saturation pulses (spoilers) applied at sufficient amplitude and duration could eliminate this artifact. However, this would not work with the supplementary coil because the gradient field is very weak in the region from which the artifact emanates.
Another application of spatial saturation at or near the edge of the field of view is brain perfusion imaging. Two images, with and without spatial saturation of the carotid arteries, are acquired and subtracted. Poor spatial saturation results in under-estimation of brain perfusion.
Yet another application of spatial saturation is 3D Time of Flight angiography with very thick slabs. This technique does not typically use a spatial saturation pulse, however, the use of a superior spatial saturation pulse can reduce signal from venous flow and improve conspicuity of arterial vessels.
Another application of gradient waveforms for localization at or near the edge of the field of view is the use of navigator pulse sequences to acquire signals from the diaphragm when performing cardiac imaging to synchronize scanning with respiration. Poor gradient linearity results in an incorrect location and field of view for the navigator signal with resulting poor scan synchronization.
The present invention is a method and apparatus for optimizing the use of a whole-body gradient coil system and a supplemental gradient coil system in clinical imaging applications where image data is acquired from a region of interest within the optimal range of the supplemental gradient coil system, and additional gradients are required to affect spin magnetization outside the region of interest. More particularly, the supplemental gradient coil system is operated during a series of imaging pulse sequences in which NMR signals are acquired from the region of interest, and gradient pulses produced by the whole-body gradient system are interleaved with the imaging pulse sequences to affect spin magnetization outside the region of interest. The interleaved gradient pulses may, for example, be employed as part of a spatial presaturation sequence or as part of a navigator pulse sequence in which transverse spin magnetization is produced in specific locations outside the region of interest.